The present disclosure generally relates to systems, devices, and methods for dispensing of therapeutic fluid and, in particular, to a device with a flow sensor or flow detector as well as to methods that makes use of such sensors or flow detectors.
Medical treatment of several illnesses requires continuous or periodic drug infusion into various body compartments through subcutaneous and intra-venous injections. Diabetes mellitus (DM) patients, for example, require the administration of varying amounts of insulin throughout the day to control the patients' glucose levels. In recent years, ambulatory portable Continuous Subcutaneous Insulin Infusion (hereinafter “CSII”) pumps have emerged as a superior alternative to the use of multiple daily injections (hereinafter “MDI”) of insulin, initially for Type 1 diabetes patients, and subsequently for Type 2 diabetes patients. These pumps, which deliver insulin at a continuous or periodic basal rate, as well as in bolus volumes, were developed to free patients from repeated self-administered injections, and to allow them to maintain a near-normal daily routine. Both basal and bolus volumes have to be delivered in precise doses, based on individual prescription, because an overdose or under-dose of insulin could be fatal. In the context of the present disclosure, “continuous delivery” includes a quasi-continuous delivery where small drug amounts are delivered in time intervals of typically some minutes, resulting in the pharmacological effect being substantially identical to a steady continuous delivery.
Insulin administration of basal and bolus doses is dependent on body glucose levels. Diabetes patients generally monitor their glucose levels and adjust insulin dosing accordingly. Glucose levels may be monitored by using blood-sensitive test strips (obtaining a blood sample through finger pricking), or by using removable insertable subcutaneous sensors. Insulin pumps can receive glucose measurements from glucose monitors either manually (inputting numbers with a keypad) or by automatically communicating (e.g., wirelessly) glucose readings from a remote glucose monitor.
Some portable infusion pumps include “pager-like” devices, where such a pager-like device includes a reservoir contained within the device housing. These devices are provided with a tube for delivering insulin from the pump which is, for example, attached to a patient's belt to a remote insertion site. The tubing length is in a range of typically 30 cm to 1.5 m. Pumping is achieved, for example, by linear movement of a piston/plunger within a reservoir in a syringe-like way, forcing fluid to be expelled from the reservoir to the outlet port. A processor-controlled motor and gear arrangement provides controlled rotational motion that is converted to linear movement by a rotation of a nut over a plunger rod drive screw.
Both basal and bolus volumes delivered in these “pager-like” devices are typically controlled via a set of buttons provided on the device. A user interface screen is typically provided on the device housing to provide the user with information about fluid delivery status, to program flow delivery, and to provide alerts and alarms. These devices represent a significant improvement over a regiment based on multiple daily injections, but, nevertheless, suffer from several drawbacks, among which are the large size and weight of such devices, their long delivery tubing, and lack of discreetness.
To avoid the consequences of comparatively long tubing for connecting pump and cannula, a new concept, an alternative architecture was proposed. This architecture is based on a remote controlled skin adherable device with a housing having a bottom surface adapted for contact with the patient's skin, a reservoir disposed within the housing, and an injection needle in communication with the reservoir. In these devices, the user interface is provided as a separate remote control unit that contains operating buttons and a screen to provide fluid delivery status, to program flow delivery, to provide alerts and alarms, and the like. Corresponding devices still have several limitations, including their heavy and bulky configuration, and the relative high cost resulting from their use due to the fact that the devices have to be replaced after several days (e.g., 2-3 days). Another drawback associated with this type of skin adherable devices relates to the required remote control. The user is generally totally dependent on the remote control unit and cannot initiate bolus delivery or operate the device if the remote control unit is not at hand, is lost or malfunctions.
A general limitation of current insulin pump devices is their lack of flow feedback. There is typically no monitoring or supervision of insulin flow within the delivery path after insulin is expelled from the reservoir outlet port by, for example, a plunger/piston linear movement, resulting in typical defects and/or hazardous situations, such as occlusions, air bubbles, or leakage being detected only with a large time delay of typically several hours. In dependence of the specific design, some of these situations may not be detected at all by the device.
In addition, the required insulin volume (dose) administration is achieved by programmed timing of motor operation, and counting motor or gear revolution with an encoder. This revolution is converted to a proportional linear movement of the drive screw and plunger (motor gear revolution reduction ratio is a fixed number). The distance of plunger/piston linear movement may be derived from the motor and gear number of revolutions and pitch of the drive screw and is proportional to the administered volume according to the drug reservoir cross section. Delivery accuracy is accordingly dependent on the precision of gears' cogwheels and drive screw pitch accuracies as well as the accuracy of the reservoir cross section. Slight deviations of those influence factors, resulting, e.g. from manufacturing and/or assembly tolerances as well as from operation wear-and-tear can affect the precision of linear movement, and consequently affect delivery accuracy. Furthermore, failures of the motor revolution counter (encoder) can cause uncontrolled motor operation, and consequently cause over- or under insulin delivery. In some cases, a long time delay for detecting defects or hazards or an over-or under delivery may result in serious medical complications, both short term and/or long/term.
A known problem of current insulin pumps both of the skin securable or pager type is the higher occurrence of severe high blood sugar events and diabetes ketoacidosis (DKA). DKA is a potentially life-threatening complication in patients with diabetes mellitus and results from an absolute or relative shortage of insulin (under or no insulin delivery). In response to glucose deprivation the body switches to burning fatty acids and producing acidic ketone bodies that cause most of the symptoms and complications. The main reasons for insulin under-delivery and consequently DKA in diabetes pump users are the occurrence of occlusion in the insulin path, air bubbles, and leakage. Occlusions occur when something blocks the infusion line. The causes can be manifold: a kink in the line, insulin crystallization, deposits of fibrin, blood clot, lipid residues, and the like. Insulin path occlusion can be detected in current pumps by monitoring pressure or torque readings from part of the insulin pump drive train (pulses generated in the processor to operate the motor). The patient is notified with an alarm when any reading exceeds a predetermined threshold.
Another implementation for detecting an occlusion in a fluid delivery tube is based on a detection of tube's radial expansion. The expansion is caused by an elevation of an upstream pressure that is caused by a downstream occlusion. Various components may be used to measure tube radial expansion, including a magnet sensitive element, resilient diaphragm, and others. In one example, an alarm is triggered by a pair of pressure sensors located at two different places along the insulin flow passage in the pump. In another example, an occlusion detector detects alteration in the shape of the insulin delivery tube.
In some of these occlusion detectors there, is a long lag time between the occurrence of occlusion and the detection of the occlusion (and alarm activation). Pressure buildup within the delivery path is usually very slow at low delivery rates typically used in insulin pump. For example, in one type of a commercial pump, occlusion is triggered by an average of 2.77 units of “missed insulin” with a typical time before alarm at a basal delivery rate of 0.05 U/h being 59.2 hours. Thus, from a practical perspective, this occlusion detector may not be able to prevent severe hyperglycemia and/or DKA, which usually occur only a few hours (e.g., 3-4 hours) after occurrence of occlusion.
Furthermore, existences of air bubbles in any medication infusion tubing can cause under-delivery. In portable ambulatory insulin pumps, especially at low programmed delivery rates, air bubbles can result in cessation of insulin administration for many hours and may consequently result in hyperglycemia and/or DKA. Tubing in currently existing insulin pager pumps extend from the pump housing to the user body insertion site, thus any air bubbles detector to detect bubbles in tubing needs to be external to the pump housing or somehow connected to the delivery tube. Typical current skin adherable insulin dispensing devices have no air bubbles detectors.
Leakage from the insulin path is another cause for insulin under delivery or completely missing delivery. Leakage can be related to cannula dislodgement from the subcutaneous insertion site. Because skin surface is usually covered by adhesive tape the user cannot see the leaking cannula. Other causes for leakage are related to leakage from the insulin delivery tube or tube connectors. Typical current insulin pumps do not have leakage detectors.
To address those drawbacks, a pump with at least two subcutaneous electrodes monitors a temporary conductivity variation in the subcutaneous tissue upon drug administration, thus allowing monitoring the correct execution of each administration.
Therefore, there is a need for a skin adherable infusion device that is inexpensive and that extends patient customization.